1. Field of the Invention
This invention pertains to non-invasive monitors, and more particularly to non-invasive monitoring of cardiac function.
2. Prior Art
Although the electrocardiogram (EKG) has been the primary non-invasive device for continuously monitoring activity of the heart in clinical medicine, it reflects solely electrical activation of cardiac muscle and provides no information on the mechanical characteristics of the cardiac pump. Consequently, the EKG may show normal or near normal waveforms in the presence of greatly impaired blood pumping capacity of the heart. Conversely, the EKG waveform may be abnormal despite normal or near normal pumping action. In terms of life support, adequate circulation of blood from the heart to the tissues, as reflected by the blood pumping capacity of the heart, is of paramount importance.
Obviously, non-invasive techniques for monitoring the blood pumping capacity of the heart are preferred over invasive ones. Nevertheless, invasive cardiac monitoring techniques, because of their perceived greater accuracy and ability to provide continuous monitoring, continue to be employed in, for example, critically ill patients. Invasive techniques generally have as their basis a catheter, such as a Swan-Ganz catheter, placed such that its tip lays within the pulmonary artery. This provides continuous recording of pressures in the pulmonary artery, and in certain instances pressures in the right ventricle, right atrium and indirectly the left atrium (pulmonary capillary wedge pressure). Injection of inert dye or cold saline from the catheter allows discrete measurements of cardiac output by dye dilution method or thermodilution, respectively. Alternatively, sampling blood for oxygen content in the pulmonary artery and a systemic artery together with measurement of oxygen consumption permits calculation of cardiac output by the Fick principle.
However, insertion of a cardiac catheter into the body may be hazardous. Its use can lead to death, which occurs in 1% of cases, and morbidity, which occurs in 33% of cases, as a result of infection and/or damage to the heart valves, cardiac arrhythmias, and pulmonary thromboembolism. Errors of technique, measurement, judgment and interpretation are common. It has been estimated that one-half million Swan-Ganz catheters used in the United States in 1986 resulted in the death of as many as 1000 or more patients. Furthermore, cardiac catheters cannot be kept in place for more than a few days owing to hazards from infection. They are also costly and laborintensive since catheterized patients require intensive care units which cost two to five times more than standard semi-private beds. In addition, health care workers face the risk of AIDS acquired virus and hepatitis virus as a result of exposure to blood of the infected patient during catheter introduction and subsequent maintenance.
Moreover, cardiac catheters do not directly provide measurement of change in ventricular volume. While such measurements can be indirectly obtained in conjunction with injection of radiopaque dye and roentgenographic imaging, this technique is time-consuming and costly, and dangerous hypotension and bradycardia may be induced by the dye. Furthermore, the number of studies in a given patient is limited by the hazards of x-ray exposure and radiopaque dye injections.
Angiographic techniques provide the most widely accepted means for measuring ventricular volumes. They allow calculation of the extent and velocity of wall shortening and of regional abnormalities of wall motion. When they are combined with measurement of pressure, both ventricular compliance and afterload (i.e., the forces acting within the wall that oppose shortening) can be determined. When the results are expressed in units corrected for muscle length or circumferences of the ventricle, comparisons can be made between individuals with widely differing heart sizes.
Cineangiography provides a large number of sequential observations per unit of time, typically 30 to 60 frames per second. Although contrast material can be injected into the pulmonary artery and left atrium, the left ventricle is outlined more clearly when dye is directly injected into the ventricular cavity. Therefore, the latter approach is used in most patients, except in those with severe aortic regurgitation in whom the contrast material may be injected into the aorta, with the resultant reflux of contrast material outlining the left ventricular cavity.
Injection of a contrast agent does not produce hemodynamic changes (except for premature beats) until approximately the sixth beat after injection. The hyperosmolarity produced by the contrast agent increases the blood volume, which begins to raise preload and heart rate within 30 seconds of the injection, an effect that may persist for as long as two hours. Therefore, this technique cannot be utilized for repetitive measurements within a short time span. Further, contrast agents also depress contractility directly, though newer nonionic agents have been found useful for minimizing these adverse effects.
In calculating ventricular volumes or dimensions from angiograms, it is essential to take into account and apply appropriate correction factors for magnification as well as distortion produced by nonparallel x-ray beams. In order to apply these correction factors, care must be taken to determine accurately the tube-to-patient and tube-to-film distances. Correction is best accomplished by filming a calibrated grid at the position of the ventricle. Thus, angiographic methods do not have wide clinical application owing to their complexity, safety considerations, invasiveness, and side effects of the contrast agents.
The importance of measuring changes of ventricular volume was well expressed by Davila in a symposium on measurement of left ventricular volume. He pointed out that the description of the functional mechanics of the left ventricle requires measurement of force, strain and velocity (rate of strain). Pressure, a standard measurement in cardiac catheterization laboratories, critical care units and operating rooms, is not necessarily dependent on shape (geometry) or size (volume) of the ventricle. However, force and strain must be expressed in relation to geometry and size of the fluid container.
In the same symposium, Chapman et al described a cineangiographic method for measuring ventricular volume. These workers also took into account the shortcomings of their method and made the following observations: "The ideal system for following change in ventricular volume is obviously one which is fully applicable to the free-living organism, which requires no injection of any sort, and which can be used repeatedly over long periods of time without danger or discomfort to the subject. Such a system, if it ever becomes available, can hardly be based on roentgenologic principles. But until some entirely different principle emerges and is applied, the roentgenologic principle is indispensable." A further requirement for an ideal system would be a minimum of physician or technician time for utilizing such technology and interpreting the results.
Because of the obvious advantage of non-invasive techniques over invasive ones, a continuing search has been made for reliable non-invasive methods of assessing cardiac performance. Such methods are needed particularly in detecting serial changes in cardiac function and in evaluating both acute and chronic effects of interventions such as drug therapy and cardiac operations. The five principal non-invasive methods for assessing cardiac performance are: systolic time intervals, M-mode and two-dimensional echocardiography, radionuclide angiography, gated computerized tomography (CT scanning), and gated magnetic resonance imaging (MRI). All but the first of these are alternatives to angiography for measurement of ventricular volumes and/or dimensions and therefore permit the non-invasive estimation of ejection phase indices. Other than in patients with obstruction to left ventricular outflow, wall stress (afterload) can be estimated from a combination of systemic arterial pressure, ventricular radius, and wall thickness. All four non-invasive imaging methods allow estimation of ventricular systolic and diastolic volumes; none, however, is satisfactory for continuous or near-continuous monitoring of critically ill patients.
Systolic time intervals have been usually obtained with the combination of an external transducer on the carotid artery in the neck to display its pulsations, a microphone over the heart to record heart sounds, and the electrocardiogram. This technique has never enjoyed wide popularity because of both technical and physiologic reasons: (1) reliable, reproducible recordings are difficult to obtain, (2) prominent internal jugular venous pulsations in the horizontal body posture may be superimposed on the carotid artery pulsations rendering interpretation of the carotid arterial waveform difficult, (3) accurate recording of heart sounds may be difficult to obtain particularly in patients with obesity or emphysema, (4) systolic time intervals are sensitive to many pharmacologic and hemodynamic influences including changes in left ventricular preload and afterload which may introduce misleading values, (5) changes in duration of systolic time intervals can be influenced by patient posture and time of day when recordings are made, (6) carotid pulse contours to calculate systolic time intervals can be difficult to interpret in patients with aortic valve disease, and (7) presence of congestive heart failure can either normalize abnormal values or make normal values abnormal.
Echocardiography involves ultrasonic imaging of ventricular wall motion to monitor cardiac function. With this technique, the dynamics of ventricular wall contraction and the internal dimensions of the cardiac chambers can be recorded. The apparatuses used for echocardiography encompass a wide variety of increasingly sophisticated and computer-aided imaging and analysis systems. The transducer placements on the chest require the services of skilled technicians and incorrect placements lead to misleading information. Furthermore, these systems are quite expensive, not readily portable, require that the patient be studied in the left lateral decubitus posture, and are not intended for continuous monitoring of critically ill patients throughout the day or during exercise.
In addition to the foregoing drawbacks, echocardiography has several inherent limitations. For example, all ultrasonic beams have a defined breadth and height comparable to the size of the crystal transducer face. Beyond its focal point, the beam's cross-sectional area enlarges in direct proportion to the distance from the transducer face. Therefore, in M-mode (single transducer) echocardiography, two laterally separated structures may appear in direct anteroposterior relationship.
Two-dimensional electrocardiographic techniques also produce distortions, which increase with increasing distance between the target and the central beam axis. In these instruments, axial resolution (1-2 mm) is superior to lateral resolution (4-5 mm). Because of the complex nature by which two-dimensional images are generated, artifacts may appear as intracardiac masses to the casual observer. Further, delineation of the endocardium of the left ventricle in its entirety is achieved only 70 to 80% of the time. Also, respiratory interference limits the ability to obtain continuous beat to beat recordings, particularly during exercise.
Attempts have also been made to determine left ventricular end-diastolic and end-systolic volumes from dimensions derived from echocardiography. These have met with variable success, depending on the patient population studied and whether m-mode or two dimensional echo techniques were employed. M-mode dimensions are used to calculate left ventricular volume through an application of the angiographic concept of the left ventricle as an ellipsoid. However, M-mode echocardiography allows measurement of only one left ventricular dimension, the septalposterolateral dimension, which is viewed at the level of the chordae tendineae. Consequently, to calculate volume from this single dimension, the following assumptions are made: (1) the ventricle being examined does in fact approximate the geometry of an ellipsoid, both in diastole and systole; (2) the septal-posterolateral dimension measured coincides with the minor axis of the ellipsoid; (3) the orthogonal minor axis is equal to the measured minor axis; and (4) the major axis is twice the length of the minor axes. While good correlations between angiographic and echo left ventricular volumes have been obtained, correlations are poor in patients who have asynergetic ventricular wall motion, which occurs in patients with coronary artery disease in whom damaged areas of the left ventricular wall do not move in phase with the normal areas. Also, because ventricular volume curves as a function of time cannot be derived without utilization of several assumptions and approximations, they are not usually reported.
Two-dimensional echocardiography offers considerable advantage for estimation of left ventricular volume because it allows direct measurement of all three hemiaxes on the ellipsoid model and also allows application of other volume formulations, such as Simpson's rule. Studies have shown that correlations between echocardiographic and angiographic volumes are substantially improved when two-dimensional methods are used, and good correlations have been obtained even in the presence of ventricular asynergy. The greatest disadvantage to quantitative two-dimensional echocardiography is the inability to obtain technically satisfactory images in all patients and the labor involved in analyzing the studies. This technique, as with the M-mode, does not readily provide dynamic changes of ventricular volume over time.
Echocardiography has also been employed to estimate the velocity of ventricular circumferential fiber shortening (Vcf). This echo measurement is analogous to the derivative of change in ventricular volume during systole and serves as a measure of ventricular contractility. Its application in M-mode echocardiography assumes that the left ventricular internal dimension is measured at the midventricular level. The mean rate of shortening is determined by dividing the calculated circumference expression by the left ventricular ejection time (ET), which may be measured from the concomitant carotid pulse tracing or from the time duration of echocardiographic aortic valve opening. Peak Vcf can be similarly derived by extrapolation from the maximum systolic slope of posterior and septal walls. Vcf is inaccurate in patients with asynergetic movement of the left ventricle as in patients with ischemic heart disease.
Mean velocity of circumferential fiber shortening (V.sub.cf) can be determined simply from measurements of end-diastolic and end-systolic dimensions by echocardiography, CT scanning, or MRI. Since the ventricle is approximately circular at its minor axis the circumference is equal to diameter (D). Mean V.sub.cf (in circumference/sec) is therefore the difference between end-diastolic and end-systolic circumference (in cm) divided by the product of the duration of ejection (in sec) and the end-diastolic circumference. Values of V.sub.cf obtained by echocardiography compare closely with those determined from cineangiograms.
Echocardiography has also been employed to estimate stroke volume (SV), which is the difference between end-diastolic volume and end-systolic volume. This technique suffers from the inherent lack of accuracy in volume estimations and, clinically, stroke volume varies widely with different physiologic circumstances such as body size, heart rate, posture and exercise. It is, therefore, not as useful a measurement as contractility. Nevertheless, provided that subjects with left ventricular asynergy are excluded from analysis, fair correlations have been reported between stroke volume derived from M-mode echocardiographic and two dimensional echo techniques on the one hand, and both thermodilution and angiographic stroke volume measurements on the other.
Another non-invasive technique is the apex cardiogram which is obtained by employing a transducer over the maximal cardiac impulse on the anterior surface of the left hemithorax in combination with the electrocardiogram. This technique is of limited usefulness for several reasons. In particular, the recording of the apex cardiogram is strongly affected by the characteristics of the recording transducer and coupling of the transducer to the skin surface. In the absence of a palpable cardiac impulse on the chest, which may occur in patients with emphysema, the apex cardiogram cannot be obtained. Moreover, interpretation of the apex cardiogram waveform for heodynamic measurements is even more problematic than systolic time intervals.
Another non-invasive device for monitoring cardiac function in the kinetocardiograph. This device records localized chest wall movements with a transducer consisting of a small metal arm attached to a flat end piece which directly contacts the chest wall. Motion of the metal arm is transmitted to a bellows, connected to a piezoelectric or strain gauge transducer.
The bellows and pickup are mounted from a crossbar over the bed, and the end piece can be placed perpendicular to any location on the chest. The amplified signal, denoted the kinetocardiogram (KCG), is obtained during breath holding at end-expiration. The KCG measures low frequency inward and outward chest movements, which range from 5 microns in the left axilla to 200 microns directly over the precordium.
Kinetocardiography differs from apex cardiography in which outward movements are accentuated by an air displacement funnel transducer placed over the apex of the heart (a position where pulsations can be felt by the examiner). For example, the KCG senses true displacements of the precordium because of its external crossbar frame of reference, whereas the apex cardiogram senses relative rib cage interspace motion. Also, the KCG is sufficiently sensitive so that records can be obtained from many points over the precordium and not just at the apex as with the apex cardiograph.
KCG recordings in humans were initially described in locations where the precordial electrocardiographic electrode leads were conventionally positioned. In these locations, the KCG generally depicts inward motion of the chest wall following the QRS wave of the electrocardiogram followed by a large number of low frequency vibrations superimposed upon an upward, outward motion. The investigators who initially described the KCG attributed the chest movements to a combination of the following factors: (1) movements due to the cardiac impact against the chest wall, (2) changes in the intrathoracic blood volume as the result of ejection or filling of the heart, (3) impact of blood in the great vessels against the chest wall and (4) positional and shape changes of the contracting and relaxing heart. Tracings of KCG over the anterior and posterior rib cage reveal: (1) a carotojugular type of pulse tracing in the infraclavicular area (attributed by the investigators to a mixed arterial venous pulse transmitted from the subclavian or axillary blood vessels), (2) with the subject prone, a waveform configuration similar posteriorly to the V.sub.4 electrocardiographic electrode placement position, and (3) with upright posture, a smaller amplitude, noisy opposite deflection signal at a posterior position corresponding to the anterior KCG signal. The investigators attributed these findings to a combination of the factors listed above.
The KCG depicts precordial outward systolic bulges in approximately 66% of patients with known myocardial infarctions. The largest outward motion is found most often at the V.sub.3 electrocardiographic electrode placement position. Outward precordial bulges occur during exercise in about 30% of patients who develop anginal pain.
Although the KCG appears to provide useful information on the mechanical properties of heart muscle, it has never received widespread clinical acceptance. This is probably because of: (1) the unwieldy transducer to patient interface; (2) restriction of patient movement and need for breathholding during recording; (3) noisy, often uninterpretable signals; (4) requirement of a great deal of skill to interpret recordings from different locations on the rib cage; and (5) lack of quantitation of the KCG waveforms with respect to changes of ventricular volume events obtained from analysis of the recordings.
Another non-invasive device for monitoring cardiac function is the cardiokymograph (CKG). This device, available from Cardiokinetics, Seattle, Wash., consists of a circular, flat capacitive plate mounted in a plastic ring strapped to the chest. Tissue motion beneath the transducer distorts an induced electromagnetic field which in turn alters the frequency of the oscillator plate. This change of frequency is converted to a change of voltage proportional to the chest wall motion at the transducer site and then displayed as an analog waveform. The CKG provides waveforms during breathholding quite similar in appearance to the kinetocardiogram. It depicts left ventricular wall motion abnormalities just like the KCG and therefore can be used to improve the diagnostic accuracy of exercise testing as an additional marker of myocardial ischemia.
The cardiokymogram suffers from the same limitations as the kinetocardiogram, namely, (1) an unwieldy transducer to patient interface; (2) restriction of patient movement and need for breathholding during recording; (3) noisy, often uninterpretable signals; (4) requirement of a great deal of skill to interpret recordings from different locations on the rib cage; and (5) lack of quantitation of the CKG waveforms with respect to changes of ventricular volume events obtained from analysis of the recordings.
Electrokymography and radarkymography are still other techniques for non-invasively monitoring cardiac function. The motions of the borders of the cardiovascular shadow obtained with roentgen rays can be visualized directly on a fluoroscope by using a photomultiplier tube to give a phasic analog signal from cyclic variations in light produced by movement of the underlying heart border (electrokymography), or from a video monitor of the fluoroscopic image and similar tracking technology (radarkymography). A graphic record of the segmental motion on the left heart border provides recordings which closely resemble the contour curve of changes in left ventricular volume over time.
Such technology can be utilized to diagnose localized segmental dysfunction of the ventricular wall. For example, radarkymography has been used to diagnose ventricular wall abnormalities, including asynergistic and akinetic motion, associated with acute myocardial infarction. Radarkymography compares favorably with left ventricular cineangiography in the diagnosis of asynergistic myocardial contraction.
However, radarkymography and electrokymography can be used only where an interface is visualized between the cardiac silhouette and adjacent structures. Poor visualization is encountered in pulmonary fibrosis, pulmonary edema, pleural fibrosis and bony distortions of the rib cage. Dyspneic patients are difficult to study since extraneous motions of the heart caused by respiration introduce artifacts. Finally, both methods subject the patient to exposure to Roentgen rays and this hazard prevents their use in situations requiring long term monitoring.
A still further non-invasive technique for monitoring cardiac function is impedance cardiography. It has long been recognized that the passage of a high frequency, low electrical current signal between electrodes placed on the heart or directed through the heart across the intact thorax produces changes of electrical impedance which varies directly with the length and inversely with the cross-sectional area of the conductor.
In impedance cardiography, detection of localized motion of the heart is highly dependent upon the placement of the electrodes. To circumvent the problems of electrode placement, the entire thorax is treated as a conductor by placing exciting and receiving electrodes at the upper and lower borders of the thorax. This permits estimation of the magnitude of cardiac stroke volume as the difference in impedance between systole and diastole. Absolute values of cardiac stroke volume (amount of blood ejected by the heart per beat) are obtained by incorporating the rate of change of impedance (an index of the velocity differences in pulse volume) into an empirically derived equation. It is the derivative waveform of torso impedance that forms the basis for its measurement by the commercial device, the Minnesota impedance cardiograph, for calculating cardiac output.
Although impedance cardiograms were initially recorded during breathholding to eliminate impedance changes superimposed by respiration, it has been found that ensemble-averaging of torso impedance waveforms using the R-wave of the electrocardiogram as a trigger pulse provides comparable waveforms during normal respiration in healthy subjects at rest and exercise and in critically ill patients.
Because changes of transthoracic electrical impedance to detect changes of cardiac volume are highly dependent on electrode placement, segmental changes of cardiac volumes and accurate reproduction of volume contours over time cannot readily be recorded with such technology. On the other hand, treating all changes of hemodynamics of the entie thorax as a single conductor appears to provide reasonable estimates of stroke volume of the heart.
It has also long been recognized that heart motion produces gas flow within the lungs, though the mechanism of this phenomenon has puzzled investigators for many years. One of the earliest researchers suggested that each heart contraction sent a volume of blood out of the thorax and the consequent negative pressure inside the affectivity rigid container caused an inflow at the mouth. Although this "aspirating" effect of the heart was subsequently well documented, the observation that the flow pulses were also present in open-chest animal preparations pointed to other mechanisms.
Cardiogenic flow pulses have been attributed to direct beating of the heart against the pulmonary parenchyma. Although artifactually induced vascular pressure pulses produce flow oscillations in the airways, these oscillations can still be seen in an airway of a lobe to which the lobar branch of the pulmonary artery has been entirely obstructed. Furthermore, injection of 25-50 ml of saline into the canine pericardial sac markedly diminishes all cardiogenic oscillations within intrapulmonary conducting airways despite the presence of normal pulmonary arterial pulsations. These observations suggest that neither pulmonary vascular pulsations nor volume changes of the heart, which should not be affected by a small pericardial effusion, were responsible for cardiogenic flow oscillations.
The heart has an irregular shape and contracts with a twisting action; this results in a forceful thrust to some parts of the adjoining lung, whereas other parts follow the inward movement of the myocardium. It is these localized transient inflations and deflations which appear to produce intrapulmonary to-and-fro flow oscillations. Pericardial fluid tends to make the external surface of the pericardial sac more spherical so that rotation or twisting of the heart no longer produces a thrust against the lung, thereby diminishing cardiogenic oscillations of the air columns.
The actual redistribution of the flow pulse among intrapulmonary airways originating from the heart depends upon relative impedance of the airways. Its magnitude depends upon the force and acceleration of the cardiac movement. However, apart from the heart movement, intrapulmonary factors must also influence the pattern and extent of transmission of the pressure impulse and the zonal volume changes that it causes. Thus, whether a zone adjacent to the heart deflates or not, giving rise to a flow pulse in the airways subtending it, depends upon its time constant. The smaller its compliance and resistance, the more likely it is to respond to the cardiogenic pressure impulse by emptying. In contrast, if the time constant is high (e.g., due to increased airway resistance), minimal emptying occurs during the time of the pressure cycle, resulting in smaller or absent flow pulses in the airways.
The preceding discussion accounts for a number of experimental observations regarding recordings of expired gas flow. Thus, although cardiogenic oscillations appear on recordings of continuous expired gas concentrations in most normal subjects, patients with emphysema may not demonstrate this phenomenon. Absence of cardiogenic oscillations has been observed in patients with bronchial asthma, with oscillations reappearing after partial relief of the bronchial obstruction. Lung disease oscillations are not seen in the trachea unless they are also present within the lobar airways.
Luisada in 1942 reviewed the historical background for the designation, "pneumocardiogram", and defined it as the recording of pressure changes which occur in the air passages of the lung as a consequence of the heart beat. He noted that graphic recordings of this phenomenon were published as early as 1861 in animals and in humans in 1876. He utilized a pressure sensing transducer from one nostril while the subject breathed normally and employed electronic filtering to eliminate the slower respiratory waves. He attributed the four positive and five negative deflections of the resulting complex waveform to the following events: 1) auricular contraction; 2) papillary muscle contraction; 3) first ventricular wave; 4) peripheral pulse; 5) second ventricular wave; 6) semilunar valve closure; 7) first diastolic wave; 8) tricuspid valve opening; and 9) second diastolic wave. He believed that the multiple waveforms present in the pneumocardiogram were due to the difference between venous inflow to, and arterial outflow from, the thorax.
Blair and Wedd in 1939 measured rib cage movements from a site below the sternum by recording pressure changes within a bellows pneumograph manufactured by the Harvard Apparatus Company. The cardiogenic oscillations recorded during breathholding were attributed by the authors to excessive outflow of blood from the chest over inflow into the chest. They calculated this volume to be 30 ml by assuming that the recording below the sternum was representative of the entire thorax.
Cardiogenic oscillations during breathholding have also been observed on analog signals from devices which display the total external movements of the respiratory system. Such oscillations were noted by Lee and Dubois in 1955 who enclosed a subject within an airtight chamber, the body plethysmograph. The subject first breathheld after inspiring air and small oscillations of pressure (calibrated as a volume) were sensed from the body plethysmograph with a sensitive pressure gauge. These oscillations were attributed to the heartbeat, but no significance was attached to the resulting complex waveforms by Lee and Dubois or by the present inventor. After the recording was obtained while breathholding on air, the subject inspired nitrous oxide (N.sub.2 O), a soluble gas, which was taken up by the pulmonary capillary blood flow.
In 1961, Wasserman and Comroe modified the body plethysmographic technique of Lee and Dubois by substituting the subject's own thorax for the rigid body plethysmograph. Change in spirometric volume then reflected the exchange of gas molecules between alveoli and blood as long as thoracic volume remained constant. The latter was an important requirement of the method. Accordingly, to continuously monitor any movements of the chest or abdomen which would invalidate this requirement, two mercury in rubber strain gauges were placed around the rib cage and upper abdomen and connected together to permit analog recording of circumferential movements of the combined rib cage and abdominal compartments.
Wasserman and Comroe believed that the cardiogenic oscillations observed with their method reflected changes in thoracic blood volume. They did not consider the oscillations to be related to changes in ventricular volume. The present inventor accepted the interpretation given by Wasserman and Comroe to the cardiogenic oscillations observed with their technique and used Wasserman and Comroe's results in a review paper on measurement of cardiac output by alveolar gas exchange.
In 1965, Bosman and Lee utilized a body plethysmograph-flowmeter method "to study the effects of cardiac contraction upon changes in lung gas volumes during breathholding both with the glottis open and closed." They reported and depicted curves with multiple rises and falls from the body plethysmograph and pneumotachograph. They interpreted these complex waveforms as showing an excess of aortic outflow over venous inflow to the thorax during systole and a reverse during diastole. Using more sophisticated technology, their work confirmed the findings of Blair and Wedd.